Medical endoprostheses or implants for a wide variety of applications are known in a great variety from the prior art. Endoprostheses in the sense of the present invention are understood to be endovascular prostheses, e.g., stents, fastening elements for bones, e.g., screws, plates or nails, surgical suture material, intestinal clamps, vascular clips, prostheses in the area of hard and soft tissue as well as anchoring elements for electrodes, in particular pacemakers or defibrillators.
Stents are used today as implants in especially large numbers to treat stenoses (vasoconstrictions). They have a tubular basic mesh or hollow cylinder, which is open at both longitudinal ends. The tubular basic mesh of such an endoprosthesis is inserted into the vessel to be treated and serves to support the vessel there. Stents have become established for treatment of vascular diseases in particular. Constricted regions in the vessels can be widened through the use of stents, resulting in a larger lumen. Through the use of stents or other implants, an optimum vascular cross section, which is the first requirement for successful treatment, can be achieved, but the permanent presence of such a foreign body initiates a cascade of microbiological processes, which can lead to gradual overgrowth of the stent and, in the worst case, to vascular occlusion. One approach to solving this problem is to manufacture the stent and/or other implants from a biodegradable material.
“Biodegradation” is understood to refer to hydrolytic, enzymatic and other metabolic degradation processes in a living body, where these processes are triggered mainly by body fluids coming in contact with the biodegradable material of the implant and lead to a gradual dissolution of the structures of the implant containing the biodegradable material. Through this process, the implant loses its mechanical integrity at a certain point in time. The term “biocorrosion” is often used as synonymous with “biodegradation.” The term “bioresorption” comprises the subsequent resorption of the degradation products by a living body.
Materials suitable for the basic mesh of biodegradable implants may contain polymers or metals, for example. The basic mesh may consist of several of these materials. The common feature shared by all these materials is their biodegradability. Examples of suitable polymer compounds include polymers from the group of cellulose, collagen, albumin, casein, polysaccharides (PSAC), polylactide (PLA), poly-L-lactide (PLLA), polyglycol (PGA), poly-D,L-lactide-co-glycolide (PDLLA-PGA), polyhydroxybutyric acid (PHB), polyhydroxyvaleric acid (PHV), poly/alkyl carbonates, polyorthoesters, polyethylene terephthalate (PET), polymalonic acid (PML), polyanhydrides, polyphosphazenes, polyamino acids and their copolymers as well as hyaluronic acid. Depending on the desired properties, the polymers may be used in pure form, in derivatized form, in the form of blends or copolymers. Metallic biodegradable materials are based primarily on alloys of magnesium and iron. The present invention preferably relates to implants whose biodegradable material contains at least partially a metal, preferably magnesium or a magnesium alloy.
There are already known stents which have coatings with various functions. Such coatings serve, for example, to release medications, to arrange an X-ray marker or to protect the underlying structures.
In the implementation of biodegradable implants, the degradability should be controlled according to the desired treatment and/or application of the respective implant (coronary, intracranial, renal, etc.). For many therapeutic applications, for example, an important target corridor is for the implant to lose its integrity over a period of four weeks to six months. The term “integrity” here is understood to be mechanical integrity, i.e., the property whereby the implant has hardly any mechanical losses in comparison with the undegraded implant. This means that the implant has such high mechanical stability that the collapse pressure, for example, drops only slightly, i.e., at most to 80% of the nominal value. Thus if it has its integrity, the implant can fulfill its main function, to maintain the patency of the blood vessel. Alternatively, integrity may be defined as such a great mechanical stability of an implant that it is hardly subject to any geometric changes in its load state in the vessel, e.g., it does not collapse to any mentionable extent, i.e., it still has at least 80% of the dilated diameter under load or, in the case of a stent, there are hardly any broken load-bearing struts.
Biodegradable implants containing magnesium or a magnesium alloy, in particular so-called magnesium stents, have proven to be especially promising for the target corridor of degradation mentioned above, but they lose their mechanical integrity and/or supporting effect too soon on the one hand, while on the other hand, the decline in integrity varies greatly in vitro and in vivo. This means that with magnesium stents, the collapse pressure declines too rapidly over time and/or there is too much variability in this decline so it is too indefinable.
Various mechanisms of controlling the degradation of magnesium implants have already been described in the prior art. These are based on organic and inorganic protective layers, for example, or combinations thereof, which present some resistance to the human corrosion medium and the corrosion processes taking place there. Approaches known in the past have been characterized in that barrier layer effects are achieved, based on the best spatial separation between the corrosion medium and the metallic material, in particular the metallic magnesium, with no defects, if possible. Protection from degradation is thus ensured by protective layers having various compositions and by defined geometric spacings (diffusion barriers) between the corrosion medium and the magnesium base material. Other approaches are based on alloy constituents of the biodegradable material of the implant body, which influence the corrosion process by shifting its position in the electrochemical voltage series. Other approaches in the field of controlled degradation produce intended breaking effects by applying physical changes (e.g., local changes in cross section) and/or chemical changes to the stent surface (e.g., multiple layers having different chemical compositions locally). However, with the approaches mentioned so far, it has not usually been possible to arrange for the dissolution, which occurs due to the degradation process, and the resulting breakage of the webs to occur within the required window of time. The result is that degradation occurs either too soon or too late and/or there is excessive variability in the degradation of the implant.
Another problem associated with passivation coatings is based on the fact that stents or other implants may usually assume two states, namely a compressed state with a small diameter and an expanded state with a larger diameter. In the compressed state, the implant can be inserted by means of a catheter into the vessel to be supported and positioned at the location to be treated. At the site of treatment, the implant is then dilated by means of a balloon catheter and/or converted to the expanded state (when a shape memory alloy is used as the implant material) by heating it to a temperature above the transition temperature. Because of this change in diameter, the implant body is subjected to mechanical stress. Additional mechanical stresses on the implant may occur during production or movement of the implant in or with the blood vessel into which the implant has been inserted. With the aforementioned coatings, this yields in particular the disadvantage that the coating may crack during deformation of the implant (e.g., forming microcracks) or may be partially removed. This may cause an unspecified local degradation. Furthermore, the onset and rate of degradation depend on the size and distribution of the microcracks formed during deformation, which, being defects, are difficult to control. This leads to a great scattering in the degradation times.
Document DE 10 2006 060 501 discloses a method of manufacturing a corrosion-inhibiting coating on an implant made of a biocorrodible magnesium alloy and an implant obtainable by this method, in which, after the implant is prepared, the implant surface is treated with an aqueous or alcoholic conversion solution containing one or more ions selected from the group of K−, Na+, NH4+, Ca2+, Mg2+, Zn2+, Ti4−, Zr4+, Ce3+, Ce4+, PO33−, PO43−, HPO42−, H2PO4−, OH−, BO33−, B4O73−, SiO32−, MnO42−, MnO4−, VO3−, WO42−, MoO42−, TiO32−, Se2−, ZrO32− and NbO4−, where the concentration of the ion(s) is in the range of 10−2 mol/L to 2 mol/L. The treatment of the implant surface with said conversion solution causes anodic oxidation of the implant. This is performed either without the use of an external current source (externally currentless) or with a current source. However, neither the examples of methods described in this publication nor the electrolyte compositions completely meet expectations with regard to degradation behavior and dilatability without destruction of the layer being used for a magnesium stent.